The present invention relates to micropumps, particularly to such devices for the precise control of fluid discharge in medical and analytical applications and to a method of applying silicon micromachining technology to the manufacture of such micropumps.
Referring to FIG. 33, micropumps of the prior art include a glass substrate or base plate 200, glass plate 205 and silicon membrane 201 bonded therebetween. Membrane 201 includes diaphragm 204 formed between valves 202, 203. The diaphragm is adapted to be driven through air layer 206 by exothermic resistor 207. Glass substrate 200 has inlet port 208 and outlet port 209 in communication with a respective valve 202, 203. When air in layer 206 is thermally expanded, diaphragm 204 is displaced downwardly thereby increasing pressure within pump chamber 210. Such pressure closes inlet valve 202 and simultaneously opens outlet valve 203 thereby discharging fluid in pump chamber 210 to outlet port 209. When air layer 206 contracts, diaphragm 204 is displaced upwardly, valves 202, 203 function in reverse so that inlet port 208 draws fluid into chamber 210 and outlet port 209 prevents fluid discharge. Such micropumps precisely control the flow and discharge of minute volumes of fluid and are particularly adapted to medical and analysis applications.
Prior art methods for the manufacture of micromechanical devices including micropumps employ semiconductor etching technology including aerotropic etching and similar machining methods for forming complicated three-dimensional silicon construction. Such methods for making various shapes of joined substrates include substrate joining technology, an anode joining method for connecting substrates of glass and silicon. Silicon-formed pressure sensors have been developed as micromechanical devices, however, no established and reliable method for the manufacture of high performance micropumps is presently available.
Conventional micropumps have several shortcomings. The first of these relates to discharge performance. Two-valve type micropumps, as illustrated in FIG. 33, are easier to manufacture than three-valve type micropumps. However, two-valve devices experience gradual lowering of fluid flow volume due to pressure-differential P between the inlet and outlet ports, thereby deteriorating micropump efficiency. The characteristics curves of FIG. 34 show that two-valve type micropump flow volume Q decreases linearly with increasing pressure differential P, as illustrated by line A. In the case of three-valve type micropumps, depicted by line B, flow volume Q is substantially independent of variations in pressure differential P. The third valve provided in the flow route between the inlet and outlet valves prevents back-flow due to pressure differential P and thereby sustains constant flow volume. In the case of two-valve micropumps, however, the pressure differential P is applied directly to the outlet valve so the outlet valve experiences substantial force in the closing direction. When the two-valve type micropump is employed to administer insulin, for example, back pressure of about 600 mm H.sub.2 O prevents discharge. In medical applications, fluid discharge is generally required at a substantially constant rate until back pressure reaches about 400 mmH.sub.2 O.
The second problem relates to securing the micropump drive means. Conventional methods often result in incomplete and ineffective installation of micropump drive apparatus. Generally, piezoelectric elements have been employed as drive apparatus because of their preferable controllability. The piezoelectric element must be uniformly bonded to the micropump drive diaphragm. Bonding a very thin membrane to the diaphragm is problematic. Because the diaphragm is secured at its periphery and is apt to bend or flex during bonding. As a result, a poor bond is attained between the piezoelectric element and diaphragm. If the piezoelectric element is pressed excessively to the diaphragm, the diaphragm periphery is stretched and subject to damage. This pressing force is difficult to control resulting in a difficult bonding operation.
The third problem concerns maintaining constant discharge performance, particularly in medical applications involving, for example, the administration of insulin. Medication over-dosing is a dangerous problem. It is necessary to immediately detect injection malfunctions due, for example, to breakdown of the micropump drive apparatus, blockage in an output needle, blending air in the pump, or valve breakage. No such detection means is provided in conventional micropumps.
The method of manufacturing micropumps having a thin membrane single crystal silicon diaphragm and a valve membrane integrally formed of single crystal silicon, and glass substrates sandwiching the thin membrane diaphragm employ fabrication methods for constructing silicon pressure sensors. However, the valve membrane includes a narrow zone defining a valve portion, through which zone the glass substrates come into contact with each other creating a small gap, deteriorating the seal. It is necessary to apply pre-pressure to such construction, however there has been no suitable means of providing such pre-pressure. When the glass plate and the main thin membrane are adhered by anode joining, the valve portion adhers to the glass plate rendering the valve body useless.